Self-calibrating ct detectors, systems and methods for self-calibration

ABSTRACT

The present approach relates to self-calibration of CT detectors based on detected misalignment of the detector and X-ray source. The present approach make the detector more robust to changes against temperature and focal spot movements. The diagnostic image generated by energy resolving calibrated response signals is able to present enhanced features compared to conventional CT based diagnostic images.

CROSS-REFERENCE TO RELATED APPLICATION

This application claims priority to and the benefit of ProvisionalApplication No. 62/438,494, entitled “SELF-CALIBRATING CT DETECTORS,SYSTEMS AND METHODS FOR SELF-CALIBRATION”, filed Dec. 23, 2016, which isherein incorporated by reference in its entirety.

BACKGROUND

Embodiments of the present specification relate generally to computedtomography (CT), and more particularly to self-calibrating CT detectors,systems and methods for self-calibration.

In an imaging system, such as a computed tomography (CT) imaging system,a fan shaped X-ray beam is emitted towards an object such as a patientor a piece of luggage to image a region of interest in the object. Thebeam is typically attenuated by the object. Subsequently, the attenuatedbeam is incident on a CT detector having an array of detector elements.In response to the attenuated beam, the detector elements of the arraygenerate respective electrical signals representative of internalinformation of the object. These electrical signals are processed by adata processing unit to generate an image representative of the regionof interest in the object.

Typically, the array of detector elements is constructed to have astandard response for all detector elements. However, there may bevariations among the responses of the detector elements. Numerous otherfactors such as, but not limited to, geometric parameters of the imagingsystem, detector gain, shading effects of auxiliary components, such asanti-scatter grids or collimators, of the detectors, may influence theresponse from the detector elements in different ways. In particular,during usage, and/or over time, detector response may change due tovariations in temperature, tube spectrum, and gantry movement. It isdesirable to calibrate the CT detector to at least partly compensate fordegradation of image quality caused due to variation in detectorresponse.

Typically, a calibration is performed before the shipment of theequipment, at the time of installation of the imaging system by the enduser (e.g., hospital radiology staff) or by a field engineer employed bythe system manufacturer. Calibration values for some or all of thedetector elements are computed from the calibration. Subsequently, thecalibration values are applied to the electrical signals generated bythe detector elements during operation of the imaging equipment. Thecalibration may be repeated periodically and/or during imaging systemmodification or maintenance events to regenerate the calibration values.

However, in current systems, there are no mechanisms to detect andcorrect the variations that occur “during” the operation. This, combinedwith the fact that calibrations are often not performed as often as maybe needed (due to the time consuming nature of existing calibrationtechniques) results in less effective calibration values continuing tobe used for periods longer than may be desirable.

BRIEF DESCRIPTION

Certain embodiments commensurate in scope with the originally claimedsubject matter are summarized below. These embodiments are not intendedto limit the scope of the claimed subject matter, but rather theseembodiments are intended only to provide a brief summary of possibleembodiments. Indeed, the invention may encompass a variety of forms thatmay be similar to or different from the embodiments set forth below.

In one implementation, a method is provided for calibrating an energyresolving computed tomography (CT) system comprising an X-ray source anda pixelated detector. In accordance with aspects of this implementation,for each respective segment of a detector element comprising a pluralityof segments, a respective response signal is acquired. Each responsesignal comprises a plurality of photon counts each corresponding to adifferent energy bin of the respective segment. For the detectorelement, a first photon count for an energy bin of a first segment and asecond photon count for the same energy bin of a second segment aredetermined. The first segment and the second segment are verticallyoffset within the respective detector element. A photon count ratio isdetermined based on the first photon count and the second photon count.A misalignment angle of the detector element with reference to the X-raysource is determined based on the photon count ratio. A plurality ofgain factors are determined for the plurality of segments of thedetector element based on the photon count ratio and the misalignmentangle. A corrected spectrum for the detector element is determined usingthe plurality of gain factors and the response signals.

In a further implementation, a computed tomography (CT) imaging systemis provided. In accordance with aspects of this implementation, the CTimaging system includes a radiation source configured to emit radiationand a pixelated detector configured to generate signals in response tothe emitted radiation. The pixelated detector comprises a plurality ofdetector elements, each detector element comprising a plurality ofsegments offset in the direction of radiation propagation. The CTimaging system further includes a readout channel for each segment ofeach detector element. Each readout channel, during operation, generatesa photon count for each bin in a plurality of energy bins. The CTimaging system further includes a calibration sub-system configured to,based upon the photon counts for a first segment and a second segment ofa respective detector element: determine a first photon count for anenergy bin of the first segment and a second photon count for the sameenergy bin of the second segment, wherein the first segment and thesecond segment are vertically offset within the respective detectorelement; determine a photon count ratio based on the first photon countand the second photon count; determine a misalignment angle of thedetector element with reference to the radiation source based on thephoton count ratio; determine a plurality of gain factors for to theplurality of segments of the detector element based on the photon countratio and the misalignment angle; and determine a corrected spectrum forthe detector element using the plurality of gain factors and theresponse signals. The CT imaging system further includes an imagereconstruction unit configured to generate an output image using therespective corrected spectra of some or all of the detector elements.

In an additional implementation, one or more non-transitorycomputer-readable media storing processor-executable instructions areprovided. In accordance with aspects of this implementation, theinstructions, when executed by one or more processors, cause the one ormore processors to performs acts comprising: for each respective segmentof a detector element comprising a plurality of segments, acquire arespective response signal, wherein each response signal comprises aplurality of photon counts each corresponding to a different energy binof the respective segment; for the detector element, determine a firstphoton count for an energy bin of a first segment and a second photoncount for the same energy bin of a second segment, wherein the firstsegment and the second segment are vertically offset within therespective detector element; determine a photon count ratio based on thefirst photon count and the second photon count; determine a misalignmentangle of the detector element with reference to an X-ray source based onthe photon count ratio; determine a plurality of gain factors for theplurality of segments of the detector element based on the photon countratio and the misalignment angle; and determine a corrected spectrum forthe detector element using the plurality of gain factors and theresponse signals.

DRAWINGS

These and other features and aspects of embodiments of the presentinvention will become better understood when the following detaileddescription is read with reference to the accompanying drawings in whichlike characters represent like parts throughout the drawings, wherein:

FIGS. 1A and 1B are block diagram representations of a self-calibratingcomputed tomography (CT) system illustrating an aligned geometry (FIG.1A) and a mis-aligned geometry (FIG. 1B), in accordance with aspects ofthe present disclosure;

FIG. 2 depicts a side-sectional view of a segmented pixelated detector,in accordance with aspects of the present disclosure;

FIG. 3 depicts the readout chain of a detector, in accordance withaspects of the present disclosure;

FIG. 4 depicts a readout channel of a segment of a detector element, inaccordance with aspects of the present disclosure;

FIG. 5 depicts an aligned detector element, in accordance with aspectsof the present disclosure;

FIG. 6 depicts a misaligned detector element, in accordance with aspectsof the present disclosure;

FIG. 7 depicts a detector element sampled at offset segments andcorresponding spectrum for each segment, in accordance with aspects ofthe present disclosure;

FIG. 8 depicts a misaligned detector element, in accordance with aspectsof the present disclosure;

FIG. 9 depicts a graphical representation of gain in view of detectorelement tilt angle, in accordance with aspects of the presentdisclosure;

FIG. 10 depicts respective aligned and misaligned detector elements, inaccordance with aspects of the present disclosure; and

FIG. 11 is a flow chart of a method for self-calibration of computedtomography (CT) image, in accordance with aspects of the presentdisclosure.

DETAILED DESCRIPTION

One or more specific embodiments will be described below. In an effortto provide a concise description of these embodiments, all features ofan actual implementation may not be described in the specification. Itshould be appreciated that in the development of any such actualimplementation, as in any engineering or design project, numerousimplementation-specific decisions must be made to achieve thedevelopers' specific goals, such as compliance with system-related andbusiness-related constraints, which may vary from one implementation toanother. Moreover, it should be appreciated that such a developmenteffort might be complex and time consuming, but would nevertheless be aroutine undertaking of design, fabrication, and manufacture for those ofordinary skill having the benefit of this disclosure.

When introducing elements of various embodiments of the presentinvention, the articles “a,” “an,” “the,” and “said” are intended tomean that there are one or more of the elements. The terms “comprising,”“including,” and “having” are intended to be inclusive and mean thatthere may be additional elements other than the listed elements.Furthermore, any numerical examples in the following discussion areintended to be non-limiting, and thus additional numerical values,ranges, and percentages are within the scope of the disclosedembodiments.

While aspects of the following discussion may be provided in the contextof medical imaging, it should be appreciated that the present techniquesare not limited to such medical contexts. Indeed, the provision ofexamples and explanations in such a medical context is only tofacilitate explanation by providing instances of real-worldimplementations and applications. However, the present approaches mayalso be utilized in other contexts, such as tomographic imagereconstruction for industrial Computed Tomography (CT) used innon-destructive inspection of manufactured parts or goods (i.e., qualitycontrol or quality review applications), and/or the non-invasiveinspection of packages, boxes, luggage, and so forth (i.e., security orscreening applications). In general, the present approaches may beuseful in any imaging or screening context or image processing fieldwhere a set or type of acquired data undergoes a reconstruction processto generate an image or volume.

Embodiments of the present disclosure relate to self-calibratingcomputed tomography (CT) detectors, self-calibrating CT systems, andmethods for self-calibration of a detector. In particular, the systemsand methods disclosed herein facilitate at least partial correction ofmisalignment of an X-ray beam with respect to a detector element in anenergy-resolving, photon-counting CT imaging system.

Further, in current systems there are typically no mechanisms to detectand correct the variations that occur “during” operation. The presentapproach allows detection of misalignment of a detector element (i.e.,pixel) in real-time with respect to the X-ray focal spot. The approachfurther includes algorithmic correction of the error in the outputspectrum caused by the misalignment.

As discussed herein, in certain implementations, the energy resolving CTimaging system includes a detector having vertically segmented detectorelements from which signals are readout from different depths of arespective segmented detector element. In operation, a self-calibratingCT detector receives a plurality of response signals from the segmenteddetector elements. As discussed herein, these signals are used todetermine a misalignment angle of a detector element (i.e., a pixel)with respect to the X-ray source. In one implementation, thisinformation is used to determine gain factors to correct the errors in aspectral response caused by the misalignment and subsequently thecorrected spectrum of the pixel.

More generally, and as discussed herein, a self-calibrating CT detectoris used to generate calibration values that can be used to adjust one ormore operating parameters of a CT imaging system to generate a qualityimaging signal and/or to process the imaging signal generated by the CTimaging system to correct or enhance the signal quality. As used herein,the terms “pixel”, and “segmented detector” are used equivalently todenote detector elements of a pixelated detector. As used herein, theterm “gain factor” refers to a gain value that is applied to a responsesignal of a segmented detector. Further, as used herein, the term“scattering signal” refers to a component of a detector signalrepresentative of X-ray photons that have undergone a directional changein their flight between the X-ray source and detector. Furthermore, theterm “pile-up” refers to a detector response exhibiting a saturatingresponse when the density of X-ray photons increases beyond a thresholdvalue. As used here, the term “channel” is used equivalently andinterchangeably to refer to a combination of a plurality of segments andcorresponding read-out electronics of a detector, where the combinationis configured to generate a response signal. The term “response” refersto a response signal obtained from a segmented detector. In the case ofthe segmented detector, the response signal includes a plurality ofphoton counts corresponding to a plurality of energy bins, e.g.,different discrete energy ranges corresponding to the energy of arespective X-ray photon. Each of the photon count values among theplurality of photon counts is generated by a photon counter associatedwith a segment of the segmented detector. Further, the term “countratio” refers to a ratio of photon counts corresponding to two segmentsof the segmented detector. The term “normal count ratio” refers to aphoton count ratio of a segmented detector for a detector that isperfectly aligned with the direction of travel of the X-ray beam. Theterm “measured count ratio” refers to a photon count ratio of tworesponse signals corresponding to two segments of the segmenteddetector, obtained from measurements. In one embodiment, the measuredcount ratio refers to a ratio of photon counts corresponding to anenergy bin among the plurality of energy bins. The terms “angularposition” and “view angle” are used equivalently and interchangeablythroughout this application to refer to an angular orientation of aradiation source or a detector.

FIGS. 1A and 1B are block diagram representations of an imaging system100, such as a self-calibrating computed tomography (CT) system, inaccordance with aspects of the present specification. FIG. 1A depictsthe CT imaging system 100 in an aligned geometry while FIG. 1B depictsthe CT imaging system in a mis-aligned geometry. As shown in thefigures, imaging system 100 includes a radiation source 124 configuredto emit a radiation signal 126 and impinging an object 128 to generate aplurality of intensity signals 130. In one embodiment, the radiationsource 124 is an X-ray source, such as an X-ray tube. The object 128 maybe an organ of a subject to be examined or a piece of luggage to bescanned. The imaging system 100 also includes a pixelated detector 98having a plurality of detector elements in the form of a plurality ofpixels 102 a, 102 b, typically arranged in an m x n array.

In one embodiment, the pixelated detector 98 is an energy-resolvingphoton counting CT detector, such as may have improved contrast-to-noiseratio and the ability to perform K-edge imaging. Such anenergy-resolving, photon-counting detector may be fabricated usingsemiconductor materials as the active material, such as cadmiumtelluride/cadmium zinc telluride (CdTe/CZT) or silicon. In certainimplementations discussed herein, a detector 98 employing segmentedsilicon strips as the detector elements (i.e., pixels 102) is described.

By way of example, turning to the figures, each of the plurality ofpixels 102 a, 102 b includes a detector element (such as a siliconstrip) having a plurality of segments 106 disposed at a plurality ofdepths with respect to the path of the emitted X-rays with respect tothe surface of the pixels 102 facing the source 124. As used herein, thedirection traveled by the X-rays (i.e., the direction of X-raypropagation) with respect to the source-facing surface of the pixels 102may be denoted as “vertical” and/or may construed as corresponding to adepth dimension, hence such a segmented detector element may bedescribed as vertically segmented.

As may be appreciated, in the context of a pixelated detector 98 for usein CT, the detector 98 includes a plurality of such pixels 102 in twodimensions for each angular position among the multiple angularpositions at which X-rays are incident on the detector 98. The pluralityof vertical segments 106 within a pixel 102 may have different heightsand/or thicknesses and generate response signals that may be used forreducing effect of artifacts in signals generated by the pixels 102 ofthe pixelated detector 98.

The pixelated detector 98 is configured to receive the plurality ofintensity signals 130 and generates a plurality of response signals 108from the plurality of segments 106 for each pixel of the plurality ofpixels 102. The plurality of response signals 108 is representative ofimaging information from a region of interest of the object 128. Theplurality of response signals 108 constitute energy resolving spectralinformation. In one embodiment, each of the plurality of responsesignals 108 includes a plurality of photon counts corresponding to aplurality of energy bins. Each photon count among the plurality ofphoton counts is indicative of a number of X-rays having energycorresponding to a given energy bin. Hence, a response signal may begenerated for each pixel and energy bin of that respective pixel so asto provide, for a given readout interval, a number of photon counts foreach energy bin at each spatial location corresponding to a respectivepixel 102.

Thus, in this example, the pixelated detector 98 generates informationcorresponding to the pixels 102 in the form of a plurality of responsesignals 108. In the depicted example, the plurality of intensity signals130 corresponds to one angular position of the radiation source 124 anddetector 98. In a CT system, the radiation source and detector 98 rotateabout the object 128 to acquire X-ray transmission data. Thus, in areal-world implementation, the imaging system 100 is configured togenerate a plurality of intensity signals 130 corresponding to each ofthe plurality of angular positions (e.g., 360°, 180°+a fan beam angle(α), and so forth) covering an entire scanning area of interest.

The pixelated detector 98 is configured to generate an artifact freeresponse for the plurality of intensity signals 130 corresponding to aparticular angular position when the pixelated detector 102 is alignedwith the radiation source 124, as shown in FIG. 1A. However, as shown inFIG. 1B, a misalignment between the X-ray source 124 and a given pixel102 may introduce an error in the plurality of response signals 108.This is illustrated in FIGS. 1A and 1B, with FIG. 1B having a firsttransmission line 138 illustrating an unaligned X-ray transmissionrelative to a second transmission line 136 (depicted in FIG. 1B)illustrating X-ray transmission when the source 124 and detector 98 areproperly aligned (i.e., along an axis of the pixels 102). Further, theplurality of response signals 108 include a plurality of scatteringsignals generated from X-ray photons being deflected when travelingbetween the source 124 and detector 98. The plurality of responsesignals 108 may also include a plurality of corrupted signals generatedfrom pile-up response with respect to one or more pixels 102, wherebyresponse at a given pixel is saturated and not indicative of the actualX-ray incidence at the pixel.

The imaging system 100 further includes an image reconstruction unitconfigured to generate a diagnostic image 142. In one embodiment, thediagnostic image 142 is obtained using image reconstruction techniquesapplied to the plurality of calibrated signals obtained from theplurality of pixels 102. In one embodiment, the diagnostic image 142 isa calibrated CT image displayed on a display device 144 for assisting amedical practitioner.

As noted above, the pixel 102 of the detector 98 are verticallysegmented, and may in one implementation be segmented silicon strips. Byway of example, in the illustrated embodiment each pixel 102 issegmented into a plurality of segments 106, such as a first segment 106a disposed at a first depth represented by reference numeral 150 and asecond segment 106 b disposed at a second depth represented by referencenumeral 152. In one such example the first segment 106 a is configurednot to saturate (such as due to having a limited thickness that will notsaturate under imaging conditions. In this example, the non-saturatingfirst segment 106 a is designed to generate a response signal that isnot affected by pile-up artifact effects. The non-saturating response isgenerated by the first segment 106 a in response to the plurality ofintensity signals 130. In the depicted implementation, the secondsegment 106 b is also designed to be a non-saturating detector segmentat a different depth, e.g., beneath the first segment with respect tothe direction traveled by the X-ray beam.

The depicted example of an imaging system 100 further includes acalibration sub-system 104 having a signal acquisition system 112, aprocessor unit 114, an image reconstruction unit 118, a memory unit 120,and a memory storage device 116 interconnected with each other by acommunications bus 122. The imaging system 100 also includes a displaycommunicatively coupled to the image reconstruction unit 118. In oneembodiment, the calibration sub-system 104 is configured to receive theplurality of response signals 108 and generate a diagnostic image 142.The diagnostic image 142 is a calibrated image generated based on thecorrected spectrum signals corresponding to the plurality of pixels 102of the pixelated detector 98.

With the preceding in mind, FIG. 2 depicts the geometry of the detector98 as a sectional view showing vertically segmented segments 106 of eachpixel 102. In the frame of reference shown, the detector elements in theform of pixels 102 are segmented along the direction of X-raypropagation (i.e., the Z-axis) and the segments 106 at different depthsare of different thicknesses. For example, in the depicted example thesegments 106 increase in thickness as their depth in the z-directionincreases. In accordance with present approaches, the segments atdifferent depths each correspond to a different readout channel. Thus,in this example, there are four segments 106 (i.e., readout channels)for any given pixel 102. As discussed in greater detail below, eachsegment 106 of each pixel 102 may be readout by a given readout channelinto a plurality of energy ranges (i.e., energy bins) to generate aphoton count for each energy bin for a given readout interval or period.

With this in mind, and turning back to FIG. 1, in one embodiment,computing the corrected spectrum includes processing the plurality ofresponse signals 108 using one or more algorithms or routines executedon the processor unit 114. In one such approach, the processor unit 114is configured or programmed to select a first photon count 146 from afirst response signal and a second photon count 148 from a secondresponse signal generated by the same detector element (i.e., pixel102). In this example, the first photon count 146 and the second photoncount 148 correspond to a same energy range (e.g., energy bin) among aplurality of energy bins associated with their respective segments. Theprocessor unit 114 is also configured or programmed to determine aphoton count ratio based on the first photon count 146 and the secondphoton count 148. Further, the processor unit 114 is also configured orprogrammed to determine a misalignment angle of the respective detectorelement (used to generate the respective photon counts) with referenceto the X-ray source 124 based on the first photon count 146 and thesecond photon count 148. The processor unit 114 is further configured orprogrammed to determine gain factor for each segment of the respectivedetector element based on the photon count ratio and the misalignmentangle. The corrected spectrum corresponding to a respective pixel 102 iscomputed based on gain factors for the segments of the respective pixeland the plurality of response signals 108.

Turning back to FIG. 2, in one embodiment, the pixels 102 are separatedlaterally by anti-scatter Tungsten plates 160 (e.g., a 20 μm sheath oftungsten) along the X-axis. The tungsten plates 160 help to preventinternal scatter within the detector. The pixels 102 are, in oneimplementation, separated along the Y-axis pixels 102 by electricalbias. By way of example, an implementation of a pixel 102 may measure˜0.4 to 0.5 mm in the X-dimension and 0.5 mm in the Y-dimension and havean absorption length between 30 mm to 80 mm (such as approximately 30 mmor 60 mm) in the Z-dimension. In one implementation, the detector stripcorresponding to a pixel 102 may be divided into nine segments 106, withvariations in the segmentation length so that approximately uniformcount rates are expected along the depth of the detector. That is theincreased thickness of deeper segments 106 may be designed so as to haveuniform count rates at each segment 106 within a pixel (i.e., takinginto account absorption of X-ray photons by higher segments in the pixel102). In addition, as shown in FIG. 2, in some embodiments, a one- ortwo-dimensional anti-scatter collimator 162 may be provided at the X-rayfacing surface of the detector 98.

With the preceding in mind, FIG. 3 depicts the readout chain of anenergy-resolving, photon-counting detector 98 having verticallysegmented silicon strip detector elements (i.e., pixels 102) asdiscussed herein. In this example, the post-patient X-ray 130 has bothprimary and scatter components. The aggregate of the post-patient X-rays130 correspond to an input or incident spectrum 202 that passes throughthe anti-scatter grid 162 before it hits the detector 98, an element 102of which is shown in FIG. 3. The incident spectrum 202 is represented asa graph of photon flux as a function of X-ray photon energy. Asignificant fraction of the incoming X-ray photons are absorbed in thevarious segments 106 in the detector elements 102 of the detector.

The resulting detector response signals 108 for each respective segment106 of a pixel 102 are fed into or acquired by readout circuitry, hererepresented as ASIC channels 180 (i.e., Readout 1, Readout 2, and soforth), such that each segment 106 has a corresponding readout channel.As shown in FIG. 4, which depicts an example of a respective readoutchannel 180, the acquired signals 108 for a given segment 106 of a pixel102 are amplified (charge sensitive amplifier 182), conditioned (shapingfilter 184), discriminated in energy (comparators 186 with correspondingenergy bin thresholds) and digitized (energy bin counters 188) togenerate the energy resolved spectral response 190 for each segment 102,as shown by the example ASIC channel 180 of FIG. 4. Thus, an ASICchannel 180 generates a plurality of counts representative of the X-rayresponse signal, with the counts sorted into a plurality of energyranges (i.e., bins) defined by a plurality of thresholds.

As will be appreciated, spectral output for each segment 106 of arespective pixel 102 is different. By way of example, the intensity(count rate) is higher for segments 106 closer to the surface of thepixel 102 facing the X-ray source as the X-ray attenuation is anexponential function of the depth of interaction. The shape of thespectrum 180 for each segment is also different, with segments closer tothe surface having relatively larger contributions from the low-energyX-ray photons compared to the lower segments, which correspondingly havegreater contributions from high-energy X-ray photons.

Turning back to FIG. 3, the energy resolved spectral outputs 190 fromeach segment 106 of a respective pixel 102 are combined, such as atimage processor 114, to generate an energy-resolved pixel outputspectrum 192. By way of example, in one implementation, the finalspectrum 192 for a given pixel 102 is composed of the sum of spectrafrom each segment 106 of the pixel:

Final Spectrum, s(E _(j))=Σ_(i=1) ^(N) ^(segment) s _(i)(E _(j)).   (1)

where s(E_(j)) is the pixel output spectrum, s_(i)(E_(j)) is thespectrum from i-th segment.

As discussed in greater detail below, based on the energy-resolvedoutput spectrum 192 for a pixel 102, signals generated by the respectivereadout channels 180 for each segment 106 of that pixel 102 may begain-corrected or adjusted based on the derived spectral information.For example, current signal generated for a given segment 106 inresponse to the incident spectrum 202 may be adjusted by some determinedgain adjustment (e.g., by addition or subtraction) to derive again-corrected signal to generate individually calibrated responsesignals at the segment level for each pixel 102.

By way of example, an adder may be used to sum up the individualcalibrated response signals to generate a corrected spectrum for a pixel102 having a plurality of segments 106. The corrected spectrum 192includes a plurality of photon counts corresponding to the plurality ofenergy bins. Similarly, pixel values from the aggregated pixels 102 ofthe detector 92 that are corrected in this manner may be used togenerate a diagnostic image. In one embodiment, the calibrationsub-system 104 of FIG. 1 automatically determines the plurality of gainfactors or adjustments for the respective plurality of pixels 102 in areal-time manner to provide calibrated signals used to generate thediagnostic image 142.

With the preceding in mind, FIGS. 5 and 6 illustrate the geometry of avertically segmented detector element (i.e., pixel 102) with respect toan X-ray focal spot direction in more detail. As will be appreciated,when the detector 98 is perfectly aligned, as shown in FIG. 5, theprimary X-ray beam 130 is normally incident on the pixel segments 106(i.e., the X-ray beam 130 propagates in a direction parallel to orcoincident with the longitudinal axis 208 of a detector element 102).That is, when aligned, the axis 208 of the detector element 102 is at azero tilt angle θ with respect to the incident X-ray beams 130. In thisaligned geometry, the X-ray beam 130 generates the maximum response fromall the segments 106 within a given pixel 102.

Conversely, when the detector element 102 and focal spot are misaligned(i.e., have a non-zero tilt angle θ), as shown in FIG. 6, fewer X-rayphotons 130 are absorbed in the pixel segments 106 due to the shadowingeffects of both the anti-scatter grid 162 and intra-pixel tungstenseparator plates 160, with the deeper segments generatingdisproportionately fewer counts due to these shadowing effects. This isillustrated in FIG. 6 by shaded region 212, which increases in terms ofproportional area as depth within the pixel 102 increases. This unequaldecrease in counts in different segments results in the corruption ofoutput spectrum 192 of the segmented pixel 102.

With the preceding in mind, the present approach detects themisalignment of a detector element, i.e., pixel 102, with respect to anX-ray focal spot. Such detection can occur in real-time, such as duringa clinical scan operation. As discussed in greater detail below, focalspot misalignment, when detected, can be corrected or compensated for,such as via the algorithm discussed herein, to correct for the error inthe output spectrum caused by the misalignment. Thus, the presentapproach: (1) estimates the misalignment angle θ from the responses ofdifferent segments 106 in a vertically-segmented pixel 102; and/or (2)corrects an error in an output spectrum 192 caused by misalignment of avertically-segmented pixel 102 and an X-ray focal spot.

In particular, with respect to the detection of misalignment angle θusing a vertically-segmented detector element, the multiple segmentsalong the depth of X-ray interaction can be exploited to estimate themisalignment angle θ in a respective detector element (i.e., pixel 102).By way of example, and turning to FIG. 7, two vertical segments 106A,106B offset by a distance L are shown. In the depicted example, thesegments are relatively narrow, however the approach can be configuredto work for any two segments 106 in the detector element 102.

For this arrangement, for a focally aligned detector (θ=0°), as shown inFIG. 7, the spectrum 220B at narrow segment 2 (s₂(E_(j))) is related tothe spectrum 220A at segment 1 (s₁(E_(j))) by the X-ray attenuationfactor, e^(−μ(E) ^(j) ^()L) where μ(E) is the linear attenuationcoefficient of the material and L is the separation between the twosegments.

$\begin{matrix}{{s_{2}\left( E_{j} \right)} = {{{s_{1}\left( E_{j} \right)} \cdot e^{{- {\mu {(E_{j})}}}L}}\mspace{14mu} {or}}} & (2) \\{\frac{\Sigma_{j}{{s_{1}\left( E_{j} \right)} \cdot e^{{- {\mu {(E_{j})}}}L}}}{\Sigma_{j}{s_{2}\left( E_{j} \right)}} = 1} & (3)\end{matrix}$

However, turning to FIG. 8, when the detector element 102 is misalignedwith respect to the X-ray focal spot (i.e., θ≠0°) the above equality nolonger holds and there is an additional gain factor, g(θ), in therelation between the spectra 220A and 220B of the segments 106A and106B. This is due to the different amount of overlap between the X-raybeam and the detector segments (illustrated by the shaded regions 222A,222B in FIG. 8). The functional relation between the spectra 220 ismodified to be:

$\begin{matrix}{\frac{{s_{1}\left( E_{j} \right)} \cdot e^{{- {\mu {(E_{j})}}}L}}{s_{2}\left( E_{j} \right)} = {{g(\theta)} = \frac{A_{1}(\theta)}{A_{2}(\theta)}}} & (4)\end{matrix}$

With respect to g(θ), this term can be estimated from measured counts(s(E_(j))) as follows:

$\begin{matrix}{R = {{g(\theta)} = \frac{\sum\limits_{j = 1}^{N_{bins}}\; {{{s_{1}\left( E_{j} \right)} \cdot e^{{- {\mu {(E_{j})}}}L}}\Delta \; E_{j}}}{\sum\limits_{j = 1}^{N_{bins}}\; {{s_{2}\left( E_{j} \right)}\Delta \; E_{j}}}}} & (5)\end{matrix}$

where ΔE_(j) is the width of j-th energy bin. With the preceding inmind, the algorithm for estimating the misalignment angle θ mayinitially involve deriving the tilt angle response function g(θ) basedon the detector geometry:

$\begin{matrix}{{g(\theta)} = \frac{A_{1}(\theta)}{A_{2}(\theta)}} & (6)\end{matrix}$

The response function R can be computed from the respective spectrumoutputs 220A, 220B of the respective segments 106A, 106B:

$\begin{matrix}{R = {\frac{\sum\limits_{j = 1}^{N_{bins}}\; {{{s_{1}\left( E_{j} \right)} \cdot e^{{- {\mu {(E_{j})}}}L}}\Delta \; E_{j}}}{\sum\limits_{j = 1}^{N_{bins}}\; {{s_{2}\left( E_{j} \right)}\Delta \; E_{j}}} = {g(\theta)}}} & (7)\end{matrix}$

The equation R=g(θ) can be solved to extract the tilt angle, θ. This isgraphically illustrated in FIG. 9. As may be appreciated, the solutionis not unique for symmetric g(θ) as the negative tilt angle also givesthe same response. With respect to the graph of FIG. 9, this graph isrepresentative of a misalignment response of a segmented pixel 102misaligned with respect to an X-ray focal spot by a tilt angle θ. Inthis example, the graph includes an x-axis representing tilt angle θ anda y-axis representative of a ratio of counts in corresponding energybins of the selected detector segments 106A, 106B to determine themisalignment tilt angle θ.

In conjunction with the graph, a curve 230 is depicted that represents amisalignment response. In one embodiment, the curve 230 may be retrievedfrom a look-up-table, or alternatively may be calculated on-the-fly. Atilt angle, e.g., θ₁, is obtained based on the photon count ratio usingthe curve 706. The tilt angle is provided may be provided as an input tothe image reconstruction unit 118 of FIG. 1 for performing imagecalibration and generating a diagnostic image 142 based on the pixeloutputs, which may be calibrated to take into account the correctedspectral response in view of the misalignment angle.

With respect to spectrum correction, one aspect the present approachemploys an algorithm as generally discussed herein. By way of example,such an algorithm may estimate a misalignment tilt angle θ as describedabove. Gain factors g_(i)(θ), i=1, . . . N_(segment) are then computedthat represent the loss of signal in each vertical segment 106 of apixel 102 due to tilt, as shown in FIG. 10, where the segment 106C of apixel 102 is depicted in an aligned state (leftmost figure) and anmisaligned state (rightmost figure), with a corresponding loss of signalarea depicted in the misaligned state. The gain factors to address suchmisalignment may be given by:

$\begin{matrix}{{g_{i}(\theta)} = \frac{A_{i}(\theta)}{A_{i}(\theta)}} & (8)\end{matrix}$

As noted above, a correction to the estimated spectrum for each segmentis applied based upon these gain factors, such as in accordance with:

s _(i) ″=g _(i) ·s _(i)′  (9)

where s_(i)′ is the measured spectrum in segment i. Based upon thecorrected segment spectra, the corrected spectrum for a given pixel 102may be computed, such as by:

ŝ(E _(j))=Σ_(i=1) ^(N) ^(segment) s″ _(i)(E _(j))=Σ_(i=1) ^(N)^(segment) g _(i)(θ)·s′ _(i)(E _(j))   (10)

The corrected pixel spectra may then be used in image reconstruction asdiscussed herein.

With the preceding in mind, FIG. 11 is a flow chart illustrating amethod 240 for self-calibration of a CT image, in accordance withaspects of the present specification. At step 242, the method includesreceiving a plurality of response signals corresponding to a pluralityof segments 106 of a detector element 102 of the pixelated detector 98,where different response signals generated for a given detector elementcorrespond to photon counts for different energy bins. At step 244, themethod includes selecting a first photon count from a first responsesignal and a second photon count from a second response signal, wherethe first photon count and the second photon count corresponds to a sameenergy bin among the plurality of energy bins. The method may alsoinclude determining a photon count ratio based on the first photon countand the second photon count at step 246. At step 248, the method alsoincludes determining a misalignment angle of the detector element withreference to the X-ray source based on the first photon count and thesecond photon count. Further at step 260, the method includesdetermining a plurality of gain factors corresponding to the pluralityof segments based on the photon count ratio and the misalignment angle.The method also includes computing a corrected spectrum corresponding tothe detector element based on the plurality of gain factors and theplurality of response signals at step 262.

At step 264, availability of the corrected spectrum signals for all theplurality of pixels is verified. If the processing of the plurality ofpixels is completed, generation of a calibrated image is initiated.However, if the processing of one or more of the plurality of pixels isnot completed, the control is passed to step 242 for processing theremaining pixels, and the steps 242, 244, 246, 248, 260, 262 arerepeated to generate a plurality of corrected spectrum signalscorresponding to the plurality of pixels. If the processing of theplurality of pixels is completed, at step 266, the method furtherincludes generating a calibrated CT image based on the plurality ofcorrected spectrum signals. In the calibrated CT image generated usingthe steps of the method 240, the artifacts due to the detectormisalignment with the radiation signal is compensated. In oneembodiment, generating the calibrated CT image includes generating aplurality of corrected spectrum signals corresponding to each view angleamong a plurality of view angles. Further, generating the calibrated CTimage includes reconstructing a plurality of sub-images corresponding tothe plurality of view angles, wherein each of the plurality ofsub-images is re-constructed based on the plurality of correctedspectrum signals corresponding to each view angle. The calibrated CTimage is generated based on the plurality of sub-images corresponding tothe plurality of view angles.

With the preceding in mind, it may be noted that implementation of thepresent calibration algorithm may vary depending on the operationalcontext. By way of example, in contexts where it is known or expectedthat calibration does not change during a scan operation, thecalibration may be performed prior to or after the scan is performed toobtain the correction factors used in image reconstruction. Likewise, ifalignment of the detector pixels and X-ray focal spot varies withrotation speed, calibrations may be performed for different rotationspeeds (such as one set of calibration factors for each rotation speedof operational interest) and these calibration factors may be stored forsubsequent correction steps. Similarly, when alignment is known orexpected to be view (i.e., view angle) dependent, calibration factorsmay be obtained and stored for each view angle. It may also be notedthat the size of the calibration layer (i.e., segments) may differ. Inaccordance with the present approach, this may also be accounted for bythe calibration factors generated in accordance with the presentapproach. Lastly, it may be appreciated that the signal-to-noise ratio(SNR) for scan operations may be improved by averaging data from pixels102 within the same sensor slab.

Various systems and methods for self-calibration of CT detectorsdisclosed herein detect misalignment of the detector during the scan andperform calibration in real-time by signal processing techniques. Thetechnique provided here make the detector more robust to changes againsttemperature and focal spot movements. The diagnostic image generated byenergy resolving calibrated response signals is able to present enhancedfeatures compared to conventional CT based diagnostic images.Advantageously, the segmented detector employed in the self-calibrationtechnique may also be used to correct effect of pile-up and scatterartifact signals on the quality of resulting CT image.

It is to be understood that not necessarily all such objects oradvantages described above may be achieved in accordance with anyparticular embodiment. Thus, for example, those skilled in the art willrecognize that the systems and techniques described herein may beembodied or carried out in a manner that achieves or improves oneadvantage or group of advantages as taught herein without necessarilyachieving other objects or advantages as may be taught or suggestedherein.

While the technology has been described in detail in connection withonly a limited number of embodiments, it should be readily understoodthat the specification is not limited to such disclosed embodiments.Rather, the technology can be modified to incorporate any number ofvariations, alterations, substitutions or equivalent arrangements notheretofore described, but which are commensurate with the spirit andscope of the claims. Additionally, while various embodiments of thetechnology have been described, it is to be understood that aspects ofthe specification may include only some of the described embodiments.Accordingly, the specification is not to be seen as limited by theforegoing description, but is only limited by the scope of the appendedclaims.

1. A method for calibrating an energy resolving computed tomography (CT)system comprising an X-ray source and a pixelated detector, the methodcomprising: for each respective segment of a detector element comprisinga plurality of segments, acquiring a respective response signal, whereineach response signal comprises a plurality of photon counts eachcorresponding to a different energy bin of the respective segment; forthe detector element, determining a first photon count for an energy binof a first segment and a second photon count for the same energy bin ofa second segment, wherein the first segment and the second segment arevertically offset within the respective detector element; determining aphoton count ratio based on the first photon count and the second photoncount; determining a misalignment angle of the detector element withreference to the X-ray source based on the photon count ratio;determining a plurality of gain factors for the plurality of segments ofthe detector element based on the photon count ratio and themisalignment angle; and determining a corrected spectrum for thedetector element using the plurality of gain factors and the responsesignals.
 2. The method of claim 1, wherein the pixelated detectorcomprises a plurality of detector elements and wherein a respectivecorrected spectrum is calculated for some or all detector elements ofthe pixelated detector.
 3. The method of claim 1, wherein the detectorelement comprises a strip of semiconductor material segmented into theplurality of segments, wherein the segments increase in thickness asdistance from the X-ray source increases.
 4. The method of claim 1,wherein the dimensions of the detector elements are approximately 0.5mm×0.5 mm along a planar surface of the pixelated detector facing theX-ray source and between about 20 mm to about 80 mm in a direction ofX-ray propagation.
 5. The method of claim 1, wherein the correctedspectrum is determined in substantially real-time during a clinical scanoperation.
 6. The method of claim 1, wherein the gain factors aredetermined prior to a clinical scan operation and used to correct thespectra of data acquired during the clinical scan operation.
 7. Themethod of claim 1, comprising: acquiring different gain factors fordifferent rotational speeds of the X-ray source and pixelated detector;wherein determining the corrected spectrum for the detector elementsuses the gain factors corresponding to the rotational speed at whichspectral data was acquired.
 8. The method of claim 1, comprising:acquiring different gain factors for different view angles of the X-raysource and pixelated detector with respect to an imaged volume; whereindetermining the corrected spectrum for the detector elements uses thegain factors corresponding to the view angle at which spectral data wasacquired.
 9. A computed tomography (CT) imaging system, comprising: aradiation source configured to emit radiation; a pixelated detectorconfigured to generate signals in response to the emitted radiation,wherein the pixelated detector comprises a plurality of detectorelements, each detector element comprising a plurality of segmentsoffset in the direction of radiation propagation; a readout channel foreach segment of each detector element, wherein each readout channel,during operation, generates a photon count for each bin in a pluralityof energy bins; a calibration sub-system configured to, based upon thephoton counts for a first segment and a second segment of a respectivedetector element: determine a first photon count for an energy bin ofthe first segment and a second photon count for the same energy bin ofthe second segment, wherein the first segment and the second segment arevertically offset within the respective detector element; determine aphoton count ratio based on the first photon count and the second photoncount; determine a misalignment angle of the detector element withreference to the radiation source based on the photon count ratio;determine a plurality of gain factors for to the plurality of segmentsof the detector element based on the photon count ratio and themisalignment angle; and determine a corrected spectrum for the detectorelement using the plurality of gain factors and the response signals; animage reconstruction unit configured to generate an output image usingthe respective corrected spectra of some or all of the detectorelements.
 10. The CT imaging system of claim 9, wherein the radiationsource comprises an X-ray tube.
 11. The CT imaging system of claim 9,wherein each detector element comprises a strip of semiconductormaterial segmented into the plurality of segments, wherein the segmentsincrease in thickness as distance from the X-ray source increases. 12.The CT imaging system of claim 9, wherein the dimensions of the detectorelements are approximately 0.5 mm×0.5 mm along a planar surface of thepixelated detector facing the X-ay source and between about 20 mm toabout 80 mm in a direction of radiation propagation.
 13. The CT imagingsystem of claim 9, wherein the corrected spectra are stored on a memoryunit of the CT imaging system for use in imaging operations subsequentto the calibration operation used to generate the corrected spectra. 14.One or more non-transitory computer-readable media storingprocessor-executable instructions that, when executed by one or moreprocessors, cause the one or more processors to perform acts comprising:for each respective segment of a detector element comprising a pluralityof segments, acquire a respective response signal, wherein each responsesignal comprises a plurality of photon counts each corresponding to adifferent energy bin of the respective segment; for the detectorelement, determine a first photon count for an energy bin of a firstsegment and a second photon count for the same energy bin of a secondsegment, wherein the first segment and the second segment are verticallyoffset within the respective detector element; determine a photon countratio based on the first photon count and the second photon count;determine a misalignment angle of the detector element with reference toan X-ray source based on the photon count ratio; determine a pluralityof gain factors for the plurality of segments of the detector elementbased on the photon count ratio and the misalignment angle; anddetermine a corrected spectrum for the detector element using theplurality of gain factors and the response signals.
 15. The one or morenon-transitory computer-readable media of claim 14, wherein theprocessor-executable instructions, when executed, calculate correctedspectra for some or all detector elements of a pixelated detector. 16.The one or more non-transitory computer-readable media of claim 14,wherein the processor-executable instructions, when executed, determinethe corrected spectra in substantially real-time during a clinical scanoperation.
 17. The one or more non-transitory computer-readable media ofclaim 14, wherein the processor-executable instructions, when executed,determine the gain factors prior to a clinical scan operation and usethe gain factors to correct the spectra of data acquired during theclinical scan operation.
 18. The one or more non-transitorycomputer-readable media of claim 14, wherein the processor-executableinstructions, when executed: acquire different gain factors fordifferent rotational speeds of the X-ray source and pixelated detector;wherein determining the corrected spectrum for the detector elementsuses the gain factors corresponding to the rotational speed at whichspectral data was acquired.
 19. The one or more non-transitorycomputer-readable media of claim 14, wherein the processor-executableinstructions, when executed: acquire different gain factors fordifferent view angles of the X-ray source and pixelated detector withrespect to an imaged volume; wherein determining the corrected spectrumfor the detector elements uses the gain factors corresponding to theview angle at which spectral data was acquired.